Many anticancer drugs are taken up non-specifically by all types of cells, resulting in serious side-effects. Therefore, an ideal delivery carrier for an anticancer drug should be able to transport the drug specifically to cancer cells and release the drug molecules inside the cells to the site of their pharmacological activities. Polymer micelles have emerged recently as promising colloidal carriers for targeting poorly water-soluble and amphiphilic drugs as well as genes to tumour tissues. Using these micelles, drug targeting to solid cancers can be achieved passively through an enhanced permeability and retention effect because of their hyperpermeable angiogenic vasculature. Drug targeting can also be achieved by using a polymer sensitive to the surrounding temperature or pH. Moreover, active drug targeting can be realized by attaching biological signals, including antibodies, hormones, peptides and small compounds such as folic acid, which can recognize cancer cells, to the surface of nanoparticles. Compared to antibodies, hormones and peptides, folic acid is less expensive, more easily conjugated to nanoparticles, and more stable during transportation, storage and use. Unlike the other ligands listed, folate is nonimmunogenic, since folate is naturally found in the body. More importantly, folate receptor is frequently expressed on the surface of many human cancer cell types and cell uptake of folate-drug conjugates or folate-conjugated nano-carriers is based on folate receptor-mediated endocytosis. A similar strategy may be envisaged for targeting drug containing micelles to other diseased cells within the body.
Polymeric core-shell nanoparticles, whose shells are constructed from temperature-sensitive poly(N-isopropylacrylamide) (PNIPAAm) or its copolymers, have been well-studied. For example, doxorubicin (DOX)-incorporated micelles made from PNIPAAm-b-poly(butylmethacrylate) and PNIPAAm-b-poly(D,L-lactide) block copolymers have been reported. The core-shell nanoparticles were formed below the LCST (lower critical solution temperature) and drug release was slow. However, the micellar structure deformed at temperatures higher than the LCST, inducing DOX release. In addition, DOX release was regulated using temperature cycles through the LCST. It is expected that using thermally responsive core-shell nanoparticles, temporal drug delivery can be achieved by local heating and cooling. However such a system suffers from the disadvantage of not being easily accessible to deep tissues or tumours. An alternative approach to target drugs to tumour tissues is to use pH-sensitive carriers. The extracellular pH of most solid tumours in patients ranges from 5.7 to 7.8. The pH of the tumour interstitial fluid rarely declines below pH 6.5 and it is challenging to develop a system with such a narrow window of pH change. Recently core-shell nanoparticles made from poly(L-histidine)-b-poly(ethylene glycol) (PEG) were reported to be pH-sensitive. These nanoparticles were dissociated, thus releasing the enclosed drug, DOX, at pH from 7.4 to 6.8. However, micelles (core-shell nanoparticles) based on poly(L-histidine)-b-PEG are not stable at pH 7.4 and must mixed with poly(L-lactide)-b-PEG micelles to improve their stability. Also, the phase change in response to the external pH change was not as sharp as that induced by the temperature change.
U.S. patent application Ser. No. 10/865,681 reported pH-triggered thermally responsive core-shell nanoparticles self-assembled from the amphiphilic tercopolymer poly(N-isopropylacrylamide-co-N,N-dimethylacrylamide-co-10-undecenoic acid) [P(NIPAAm-co-DMAAm-co-UA)]. These micelles changed phase (from water-soluble, well dispersed in aqueous solution to water-insoluble, precipitated from aqueous solution) rapidly in response to an external pH change. These nanoparticles exhibited a pH-dependent lower critical solution temperature (LCST). In a normal physiological environment (pH 7.4), the LCST of the nanoparticles was well above the normal body temperature (37° C.) and the nanoparticles were thus well dispersed. In an acidic environment (i.e. tumour tissues, endosomes or lysosomes), however, the LCST was below 37° C., leading to the deformation and precipitation of the core-shell nanoparticles and to the eventual release of enclosed drug molecules. This application disclosed pH-triggered temperature-sensitive micelles made from a random copolymer of NIPAAm (temperature-sensitive), DMAAm (hydrophilic-to adjust the LCST of the polymer) and UA (a hydrophobic and pH-sensitive compound). These micelles possessed pH-dependent LCST, being higher than the normal body temperature under a simulated physiological condition (PBS, pH7.4) but lower than the normal body temperature at pH 6.6 or below. Therefore, the micelles were stable in a physiological environment but deformed and precipitated at pH 6.6 or below, releasing the enclosed drug molecules. These micelles may provide a good carrier for delivering anticancer drugs to tumour tissues (slightly acidic) or for intracellular drug delivery (escaping from the endosomes-a low pH environment and thus entering the cytosols). However, the polymer described in this application was a random amphiphilic copolymer, from which the micelles formed had a flexible core and a wide particle size distribution.
EP00822217 disclosed a diblock copolymer based on poly(ethylene oxide)-b-polyester, more specifically poly(ethylene oxide)-b-polylactide or polylactone. EP00844269 disclosed a diblock copolymer based on poly(ethylene oxide)-b-polyester or poly(methacrylic acid). The polymers claimed were similar to those disclosed in EP00822217. EP00852243 disclosed a diblock copolymer based on poly(ethylene oxide)-b-polyester or poly(methacrylic acid). On the end of poly(ethylene oxide) block, there was a sugar group. The polymers claimed were similar to those disclosed in EP00822217. All of the disclosed polymers were synthesized by ionic living polymerization, and none had pH- and temperature-sensitive functionalities.
US622903 disclosed a polymer-lipid conjugate. The hydrophilic polymer chains were releasably attached to liposomes via a disulfide bond, pH sensitive bond, enzymatically cleavable bond, or photochemically cleavable bond. After the release of the hydrophilic polymer chains, the hydrophobic segments of the liposomes were exposed to physiological membranes, providing the chance to fuse with the cell or liposome membrane.
US22082198 disclosed macromolecular micelles made from poly(ethylene glycol)-b-poly(amino acids). The amino acids carry amine groups or carboxyl groups, rendering the polymers chargeable for DNA or protein delivery. In particular, poly(ethylene glycol)-b-poly(lysine) and poly(ethylene glycol)-b-poly(aspartic acid) were disclosed. US22172711 disclosed a polymer-lipid conjugate, similar to that of US6224903.
US24077540 disclosed a pharmaceutical composition comprising a biologically active agent and a mucosal delivery-enhancing effective amount of a permeabilizing peptide. It was claimed that this pharmaceutical composition could be administered in combination with a membrane penetration-enhancing agent such as surfactant, mixed micelle and liposome. However, the types of micelles to be used were not disclosed.
WO00226241 disclosed lipid-comprising drug delivery complexes for gene delivery. The complexes comprised poly(ethylene glycol), lipid (e.g. DOTAP, DOPE, DOPC, DSPE, DLPE etc.), polycation (i.e. PEI) and targetable peptides.
WO04002404 disclosed pH-sensitive block copolymers for drug targeting to tumours. The block copolymers were made from poly(L-histidine), poly(ethylene glycol), poly(L-lactic acid) and/or poly(lactic acid-co-glycolic acid). However, the polymers contained no targeting component for targeting the copolymers to the tumours. US25025821 disclosed similar polymers to those of WO04002404. In particular, mixed micelles containing poly(L-histidine)-poly(ethylene glycol) block copolymer and poly(L-lactic acid)-poly(ethylene glycol) block copolymer were disclosed.